Positron-emission tomography (PET) has unique capabilities in diagnostics as it is sensitive to enhanced biological activity caused by trauma or disease. For example, it is highly effective in locating hairline bone fractures that can be difficult to identify by X-rays. However PET has operational drawbacks that preclude its widespread use for a broad range of diagnostics: it involves a large dose of radioactive tracer to the patient; it has limited resolution due to gamma ray scattering and statistics; PET cameras use expensive scintillator crystals; and it requires an extensive supporting infrastructure. As a result, PET is currently limited in use to large facilities, largely in cities in highly developed countries.
Current PET camera designs require expensive high-density scintillator crystals to provide good spatial resolution when viewed solely from the back face of the crystal array. The depth of interaction of the gamma rays in the scintillator is not measured directly, and a high density ameliorates the blurring of the interaction point due to the concomitant uncertainty. However the uncertainty in the depth-of-interaction is comparable to the depth of the crystal array, and is therefore still quite large.
Another disadvantage of the use of scintillator crystals is that the cost of arrays of expensive scintillator crystals has limited PET camera designs to those with small arrays and small geometric coverage, requiring scanning of the patient by a moveable camera, resulting in higher radiation dose and longer exposure time. The large radiation dose inherent in current PET diagnostics renders PET unsuitable for certain patient populations, as well as clinical settings, examples being that the use of PET for children is recommended only for extreme cases, and that PET cameras are found largely in urban hospitals. This significantly limits the prescription of PET as a diagnostic compared to other diagnostic tools, such as X-rays.
Time-of-flight PET (TOF-PET) ameliorates some of these drawbacks by supplying information on the position of the positron (e+)-electron (e−) annihilation by measuring the difference in times of arrival of the resulting gamma rays in the two opposing crystal arrays of the PET camera. This provides more information per event for image processing, allowing a lower dose and exposure. However, a drawback of conventional TOF-PET is that the crystals are typically read out with conventional photomultipliers (PMT's), with an interpolation to achieve spatial resolution comparable to the individual crystal size by using Anger logic. The photomultipliers occupy a substantial volume behind the crystals (away from the patient or source), adding expense and limiting the available geometries for TOF-PET cameras. The required volume typically precludes viewing the scintillator from the front face (toward the patient or source) in a clinical camera.